CHAPTER 3 Rational Design of Translational Nanocarriers · Rational Design of Translational Nanocarriers 33. ... Using a stepwise pH-controlled process, ... (PDEA) middle layer, and

CHAPTER 3

Rational Design of TranslationalNanocarriers

QIHANG SUNb, MACIEJ RADOSZb ANDYOUQING SHEN*a

aCenter for Bionanoengineering and State Key Laboratory of Chemical

Engineering, Department of Chemical and Biological Engineering, ZhejiangUniversity, Hangzhou 310027, P. R. China; bDepartment of Chemical and

Petroleum Engineering, Soft Materials Laboratory, University of Wyoming,

Laramie, WY 82071, USA

*E-mail: [email protected]

3.1 The Three Key Elements for TranslationalNanomedicine

Nanometer-sized drug carriers, including polymer–drug conjugates, dendri-

mers, liposomes, polymer micelles, and nanoparticles, have been extensively

investigated in drug delivery for cancer chemotherapy.1,2 Cancer drug deliveryis a process using nanocarriers with appropriate sizes (usually between several

nanometers and 200 nm) and stealth properties to preferentially carry drugs to

tumor tissues via the enhanced permeability and retention (EPR) effect.2

However, despite improved pharmacokinetic properties and reduced adverse

effects,1,3 currently cancer drug delivery has only achieved modest therapeutic

benefits.3,4 Thus, the design of nanocarriers with more efficient drug delivery

and thus higher therapeutic efficacy is still a pressing need.

The cancer drug delivery process can be divided into three stages, shown inFigure 3.1. Initially, the drug-loaded nanocarrier circulates in the blood

compartments, including the liver and the spleen. When passing through

RSC Polymer Chemistry Series No. 3

Functional Polymers for Nanomedicine

Edited by Youqing Shen

# The Royal Society of Chemistry 2013

Published by the Royal Society of Chemistry, www.rsc.org

32

tumor blood vessels, the carrier may fall into the pores in the blood vessel wall

and extravasate into the tumor tissue (EPR effect) (Figure 3.1A).5,6 Next, it

may further penetrate through the tumor tissue, which is nontrivial because of

the high cell density and high interstitial fluid pressure (IFP) (Figure 3.1B).7

Upon sticking to the surrounding cancer-cell membrane (Figure 3.1C), the

carrier is expected to enter the cells via one or several possible pathways, and

finally traverse the crowded intracellular structures and viscous cytosol to thetargeted subcellular sites and release the carried drug cargo.

Thus, to achieve efficient drug delivery from the i.v. injection site to the target in

the tumor cells, the nanocarrier must simultaneously meet two pairs of challenges

(Figure 3.1): (a) the nanocarrier must retain the drug very tightly, ideally without

any release, during the transport in the blood compartments and the tumor tissue,

but must be able to efficiently release the drug once reaching the intracellular

target to exert its pharmaceutical action; (b) the nanocarrier must be ‘‘slippery’’ or

‘‘stealthy’’ while in the blood compartments and in the tumor tissue until it reachesthe targeted tumor cells. The stealth in the blood compartments enables it to

effectively evade the reticuloendothelial system (RES) screening, particularly the

capture by liver and spleen for a long blood circulation time. As the blood

circulation time of the nanocarrier increases, so does its opportunity to pass the

hyperpermeable tumor blood vessel and extravasation into the tumor tissue. After

extravasating into the tumor, the nanocarrier must remain ‘‘stealthy’’ to penetrate

deep into the center region to deliver the drug. This region lacks vascular perfusion

Figure 3.1 Cancer drug delivery process: (A) transport in the circulation, (B) transportthrough the tumor tissue, and (C) transport in the tumor cell. The nanocarriermust meet two pairs of challenges — For the drug: the nanocarrier mustretain the drug very tightly during the transport in the blood compartmentsand the tumor tissue but efficiently release the drug once reaching theintracellular target; For the surface: the nanocarrier must be ‘‘very stealthy’’during in the blood compartments for a long blood circulation time andremain ‘‘stealthy’’ in penetrating the tumor tissues but must become ‘‘sticky’’or ‘‘cell binding’’ once interacting with tumor cells for efficient cellular uptake.

Rational Design of Translational Nanocarriers 33

but harbors the most aggressive and resistant cells. On reaching the targeted cells

the nanocarrier must become ‘‘sticky’’ or ‘‘cell binding’’ to interact with the tumor

cell for efficient cellular uptake. A nanocarrier capable of simultaneously

satisfying such opposite 2R2S requirements at the right time and the right place,

that is, ‘‘drug Retention in blood circulation vs. Release in tumor cells (2R)’’ and

‘‘surface Stealthy in blood circulation and tumor tissues vs. Sticky to tumor cells

(2S)’’, will deliver the drug specifically to the tumor cells, giving rise to high

therapeutic efficacy and few side effects.

While the 2R2S capability of a nanocarrier may render the resulting

nanomedicine efficacious and potentially safe for clinical translation, two other

elements, namely the feasibility of the nanocarrier materials to be proved for use

as excipients (referred to as material excipientability) and the ability to establish

scaled-up production processes for good manufacturing practice (GMP) for the

nanocarrier and its formulation with the drug (nanomedicine) (referred to as

process scale-up ability) are also indispensible for the nanomedicine to be truly

translational from the benchtop to the bedside (Figure 3.2).8 Most of our current

research is focused on using new material design and chemistry to improve the

2R2S capability; however, research aimed at translational applications should

comprehensively consider the other two elements at an early stage.

Herein, we briefly review the approaches addressing nanocarriers’ 2R2S

capability and summarize the factors affecting material excipientability and

process scale-up ability, aimed at promoting the developments of truly

translational nanomedicine for cancer drug delivery.

Figure 3.2 The three elements for translational nanomedicine: the nanocarriershould have the 2R2S capability and its material should be suitable forexcipient use (referred to as material excipientability); the production ofthe nanocarrier and its formulation with drug (nanomedicine) should beable to scale up for good manufacture process (GMP) (scale-up ability).Reprinted with permission from ref. 8. Copyright 2012 Elsevier.

34 Chapter 3

3.2 The 2R2S Capability of Nanocarriers

3.2.1 2R: Drug Retention in Circulation versus IntracellularRelease

3.2.1.1 Approaches to Minimize Premature Release from a StableCarrier

Figure 3.3 illustrates two examples, an ideal one for the case when the carrier

retains the drug during transport in the blood compartments and the tumor

tissue, but releases it in the tumor cells, and another for a typical case of

undesirable burst release when the carrier releases its drug cargo prematurely

while still circulating in the blood. Such a burst release is generally observed

for polymer particles9,10 and liposomes.11,12 As a result, the drug is dumped in

the blood compartments, which causes not only local or systemic toxicity, but

also lowers the drug availability to the tumor and thereby the therapeutic

efficacy.

Although the exact mechanism of burst release is still not fully understood, it

is likely that drug-diffusion resistance can help explain and control it. A study

on a model zero-order device indicated that the rate and extent of burst release

from an otherwise stable carrier were affected by drug solubility and drug

diffusion in an aqueous medium and by the drug loading content.13 Such

findings inspired more recent approaches to prevent burst release aimed at

enhancing drug loading, inhibiting drug diffusion from the carrier, or both.

Figure 3.3 Sketch of ideal controlled release vs. premature burst release. Adaptedwith permission from ref. 8. Copyright 2012 Elsevier.


(1) Using new chemical processes to fabricate structured nanoparticles.

Polymeric micelles encapsulate drugs mostly via physical trapping based on

hydrophobic interactions. They are generally fabricated by coprecipitation of the

hydrophobic drugs with the hydrophobic blocks of amphiphilic copolymers by

dialysis or the solvent-evaporation method,14 assuming that the drugs and the

hydrophobic blocks precipitate simultaneously and thus the drugs are completely

embedded in the hydrophobic micelle core. However, in many cases this is not a

very realistic assumption, as either the drugs can precipitate first or the core can

form first, which prevents proper drug encapsulation in the core. For example,

when the core forms first, most drug molecules may precipitate around the core,

which are prone to burst release upon dispersion in an aqueous solution.15

Building on this finding, we proposed that coating the core with an additional

hydrophobic layer would impose an extra diffusion barrier and thereby minimize

burst drug release. Using a stepwise pH-controlled process, three-layer onion-

structured nanoparticles (3LNPs) were synthesized that consisted of a poly(e-caprolactone) (PCL) core, a pH-responsive poly[2-(N,N-diethylamino)ethyl

methacrylate] (PDEA) middle layer, and a polyethylene glycol (PEG) outer

coronal layer.16 Compared to the conventional core-corona micelles, such 3LNPs

were found to exhibit a significantly lower burst release of camptothecin (CPT) at

physiological pH due to the effective barrier of the hydrophobic PDEA barrier.

The conventional method for preparing polymeric micelles through liquid

solvent evaporation or dialysis offers little control of micellization versus drug

precipitation. However, this can be accomplished with a near-critical fluid

micellization (NCM) method to prepare drug-loaded polymeric micelles.17 The

solvating power of a near-critical fluid solvent is easily tunable with pressure.

Thus, more selective and flexible micellization can be controlled by adjusting

the pressure alone. At high pressures, drugs and polymers were molecularly

homogenous in a near-critical solvent, whereas at moderate pressures

micellization/drug encapsulation occurred (Figure 3.4A). With this process,

PEG-PCL micelles, formed in a near-critical dimethyl ether/trifluoromethane,

could be loaded with paclitaxel (PTX) as high as 12 wt% (Figure 3.4B). More

recently, we prepared three-layered micelles formed by a stepwise NCM

process that exhibited little, if any, burst release despite the high drug loading

content (Figure 3.4C, D).18 The biggest advantage of this NCM is that it uses

the conventional Food and Drug Administration (FDA) approved materials to

obtain high drug loading micelles with minimized burst or even burst-free.

Such products are also free of contamination from organic solvents.

(2) Drug conjugation

The second approach to eliminate burst release is by conjugating drugs to the

carriers via covalent bonds. Because the drug must be released once at the target,

the covalent bonds or linkers must be cleavable in the tumor-cell environment.

For instance, doxorubicin (DOX) was conjugated to a poly(L-aspartic acid)

[P(Asp)] block in the block copolymer PEG-b-P(Asp) through amide19,20 or

hydrazone linkers.21,22 Drugs can also be conjugated to the ends of hydrophobic

blocks.23,24 The resulting micelles formed from PEG-block-poly(L-amino acid)

36 Chapter 3

(PEG-b-PLAA) conjugated drugs eliminated any burst release.25 Using labile

linkers responsive to the tumor’s extracellular or intracellular stimuli results in

drug release triggered in a tumor extracellular environment.26,27

Our group demonstrated, using drugs as the hydrophobic part, directly

making self-assembling amphiphilic prodrugs for fabricating burst-free

carriers.28 In this method, hydrophobic CPT molecules were conjugated to

short oligomer chains of ethylene glycol (OEG) to form the amphiphilic

phospholipid-mimicking prodrugs OEG-CPT or OEG-DiCPT (Figure 3.5).

These prodrugs formed stable liposome-like nanocapsules with extremely high

drug loading content but no burst release. Similar nanoparticles were prepared

from an amphiphilic curcumin prodrug.29

Figure 3.4 Nanoparticles prepared by a near-critical fluid micellization (NCM)method: (A) Micellization and cloud pressures of PEG-PCL and PTX in70% dimethyl ether/30% trifluoromethane. (B) Drug loading contents oftamoxifen and PTX in PEG-PCL nanoparticles by the NCM process andconventional solution process. Adapted with permission from ref. 17.Copyright 2011 American Chemical Society. (C) Cumulative drug releaseas a function of time for a diblock and two triblocks copolymers. Solidlines represent micelles prepared by near-critical micelliazation whiledashed lines represent micelles prepared conventionally by solventevaporation. (D) Cumulative drug release plotted as a function of t1/2

for experiments plotted in (C). Adapted with permission from ref. 18.Copyright 2012 American Chemical Society.


The main disadvantage of such conjugation approaches is that they may

change the drug chemical structure,30 which in turn may reduce the

pharmaceutical efficacy, not to mention the need for extensive preclinical

tests and clinical trials before acquiring FDA approval.

(3) Core- or shell-crosslinked micelles

The third approach aimed at reducing the burst release is crosslinking the

core or the corona shell of micelles. For example, Wooley et al. developed

methods for fabricating shell-crosslinked micelles.31 In order to de-crosslink

the shell to allow drug release at the target site, a linker labile in the presence of

intracellular glutathione (GSH) was used.32 As intended, such crosslinked

shells inhibited drug diffusion from the micelles, and hence reduced burst

release. However, such a crosslinked shell becomes more rigid and hence loses

its ability to repel serum proteins or other biomacromolecules,33 and thus may

not continue to be stealthy in circulation.

Covalent crosslinking of the micelle hydrophobic core can therefore be a

preferable approach.34,35 For instance, crosslinked micelles consisting of PEG-

b-poly(acryloyl carbonate)-b-poly(D,L-lactide) (PEG-PAC-PDLLA) had high

stability and significantly inhibited PTX release at low micelle concentrations

compared to the non-crosslinked controls.36 Lavasanifar et al. applied click

chemistry and developed hydrolysable core-crosslinked PEG-b-poly(a-propar-gyl carboxylate-e-caprolactone) (PEG-PPCL) micelles that exhibited a lower

degree of PTX burst release than equivalent non-crosslinked micelles.37 When

the crosslinked core had disulfide linkers, it was shown to hold the drug tightly

but release it quickly once in the tumor cell, due to the cleavage of the

crosslinkages by intracellular GSH.38 Similarly, thiolated Pluronic copolymer

(Plu-SH) was demonstrated to form core-crosslinked micelles that were

reversible via dithiothreitol (DTT)-breakable disulfide bonds, which inhibited

the premature release in an aqueous solution.39

Figure 3.5 Amphiphilic CPT prodrugs (OEG-CPT and OEG-DiCPT) and their self-assembly into nanocapsules. Reprinted with permission from ref. 28.Copyright 2010 American Chemical Society.

38 Chapter 3

3.2.1.2 Approaches to Increase Carrier Stability to PreventPremature Release

A thermodynamically unstable carrier (an unstable carrier for short) may

dissociate before reaching its target and thus prematurely release the drug.

Such an unstable carrier may dissociate fast or slowly, referred to as micelledissociation kinetics (some authors25 use ‘‘kinetic stability’’). We always prefer

carriers that are thermodynamically stable until they reach their target. At a

given temperature, micelles form at the polymer concentrations above the

critical micelle concentration (CMC):40

CCMC& exp {neh=kbTð Þ

where kbT is the thermal energy and eh is the monomer effective interactionenergy with the bulk solution (related to x in polymer physics). Polymers with

a low CMC suggest a high thermodynamic stability, and vice versa. Usually,

the longer the hydrophobic blocks, the more stable the micelles they form.41

Thermodynamic stability is particularly important because locally, in

circulation, micelles may dissociate if the block copolymer concentration falls

below the CMC. It seems intuitive that a drug-loaded micelle may have a CMC

that is different from its virgin drug-free analog, but, to a first approximation,

it is common to neglect this difference.Once the copolymer concentration falls below its CMC, the micelle

dissociation rate can vary, depending on cohesive forces among the core-

forming blocks. Chain insertion/expulsion and micellar fusion/splitting are two

mechanisms that can explain the overall dynamic exchange between monomers

and micelles.41 Monte Carlo simulation indicated that chain insertion/expulsion

played a major role when the polymer concentration was low.42 Because chain

mobility plays a crucial role, the hydrophobic blocks with a relatively high glass

transition temperature (Tg) make the micelles dissociate much more slowly thanthose with a low Tg.

43 Furthermore, the size of the hydrophobic block and the

hydrophilic-to-hydrophobic block mass ratio were found to affect the rate of

micelle dissociation, from size-exclusion chromatography (SEC) experiments.

For simple PEG-PCL copolymers, micelles formed from PEG-PCL (5000:4000

and 5000:2500) dissociated slowly; however, micelles formed from the PEG-PCL

(5000:1000) dissociated quickly into monomers.44

Even though there is evidence that some polymeric micelles can be stable in

serum even in vivo,45 the stability of micelles in the blood is far from understood.Quite different from carriers tested in water or in buffer solutions, micelles in

blood circulation can be extremely diluted and encounter various blood

components which may promote micelle dissociation. Burt et al. prepared

radiolabeled PTX-loaded PEG-PDLLAmicelles and found that PTX was rapidly

released from the micelles, and the diblock copolymer was cleaved into its two

polymer components in the blood.46,47 Maysinger et al. conjugated fluorescein-5-

carbonyl azide diacetate to PEG-PCLmicelles and noticed that they were stable in

buffer solutions but unstable in serum-containing culture media with or without


cells.48 Recently, Cheng et al. employed a fluorescence-resonance energy-transfer

(FRET) technique to demonstrate that PEG-b-PDLLAmicelles were not stable in

the bloodstream due to the influence of a- and b-globulins rather than c-globulinor serum albumin.49 Based on those results, Cheng et al. summarized the possible

mechanisms responsible for the micelle decomposition induced by serum

proteins,41 including protein adsorption,48,49 protein penetration,50,51 and drug

extraction.41 What exactly happens to the micelles after injection is poorly

understood because it is hard to measure and estimate micelle concentration

locally in the bloodstream.52 Cheng et al. tracked unmodified copolymer micelles

using the FRET imaging method, but unfortunately no direct evidence proved

that the CMC was unchanged by incorporating a FRET pair.53

However, there is no doubt that, directionally, the lower the CMC, the higher

the probability of micelle stability in the bloodstream. Therefore, the most

common strategy to enhance the micelle stability is to reduce its CMC. Compared

to liposomes, polymeric micelles usually have amuch lower CMC, at amicromolar

level, which imparts a higher stability. A further reduction of polymeric micelle

CMC can be achieved by increasing the core-forming block hydrophobicity,

molecular weight, or both.40 One example is that of chemically modified Pluronics:

Pluronic/PCL copolymeric nanospheres exhibited a lower CMC.54,55 Another

interesting finding is that stearic acid as side chains can keep micelles stable even in

the presence of serum.56 In the presence of serum albumin, a- and b-globulins, or cglobulins, the micelles from PEG-b-poly(N-hexyl stearate L-aspartamide) (PEG-b-

PHSA) copolymers with nine stearic acid side chains still existed after two hours.

Crosslinking is a straightforward method to stabilize micelles. While the

covalent crosslinking of the micelle core or shell can inhibit burst release from a

stable micelle, it can also inhibit or prevent micelle dissociation. For instance,

PEG-PCL micelles with cores crosslinked by radical polymerization of the

double bonds introduced to the PCL blocks turned out to be more stable.57

Biodegradable thermosensitive micelles with crosslinked cores formed from

PEG-b-[N-(2-hydroxyethyl methacrylamide)-oligolactates] [PEG-b-p(HEMAm-

Lacn)] kept their integrity upon dilution and only degraded after cleavage of the

ester bonds in the crosslinkers.58

The caveat, however, is that crosslinking reactions usually occur after core

formation, which can alter the structure and properties of the encapsulated

drug. To overcome this potential problem, our group developed stable core-

surface crosslinked micelles (SCNs), shown in Figure 3.6, made from

Figure 3.6 Formation of SCNs from amphiphilic brush polymers. Adapted withpermission from ref. 59. Copyright 2004 American Chemical Society.

40 Chapter 3

amphiphilic polymer brushes.59 The key point is that the backbones of the

polymer brushes acted as crosslinkages on the hydrophobic core surface,

instead of chemical crosslinking, which substantially enhanced the micelle

stability. Specifically, the resulting micelles had much lower CMCs than

corresponding PEG-PCL block copolymers.

For the excretion of the nanocarriers from the body, crosslinked micelles must

be able to break into small polymer chains. Toward this end, reversible

crosslinking triggered by different stimuli like pH,60 UV light,61 or others62 was

later developed. Historically, pH-sensitivity was the first one used to trigger a

desired carrier change because cancer or inflammation makes the extracellular

pH at the disease site acidic.63 For instance, micelles formed from the triblock

copolymer PEG-b-poly[N-(3-aminopropyl)methacrylamide]-b-poly(N-isopro-

pylacrylamide) (PEG-PAPMA-PNIPAM) were shell-crosslinked with ter-

ephthaldicarbaldehyde (TDA) at pH 9 via cleavable imine linkages.60

However, at pH , 6 the hydrolytic cleavage of the imine crosslinkages occurred.

Other examples64,65 were inspired by crosslinking, using disulfide linkages that

are sensitive to intracellular GSH (y0.5–10 mM as opposed to y20–40 mM in

the bloodstream66). For example, micelles made of a PCL-b-poly[(2,4-

dinitrophenyl)thioethyl ethylene phosphate]-b-PEG (PCL-b-PPEDNPT-b-PEG)

triblock copolymer, crosslinked with disulfide bonds, were found to be stable in

circulation but quickly decomposed in intracellular fluid.65

Even if the micelle happens to be unstable, its decomposition rate can be reduced

by choosing a stiff or bulky core. Toward this end, benzyl groups were introduced

to increase the rigidity of hydrophobic cores.67 Lavasanifar et al. synthesized benzyl

carboxylate-substituted e-CL monomers and prepared PEG-b-poly(a-benzylcarboxylate e-caprolactone) (PEG-b-PBCL) copolymers.67 For comparison, they

also prepared PEG-b-poly(a-carboxyl-e-caprolactone) (PEG-b-PCCL) by further

catalytic debenzylation. Their results demonstrated that the stability of micelles

with core structures containing aromatic groups (PEG-b-PBCL) was higher than

that of the parent PEG-PCL micelles and of the PEG-b-PCCL micelles. The

micelle decomposition rate can also be reduced by crystallizable hydrophobic

blocks.45,68 Another approach is to enhance ionic or hydrogen bonding interactions

in the micelle core. For example, polyion complex (PIC) micelles with oppositely

charged macromolecules, such as DNA or peptides, are resistant to enzymes in the

bloodstream,69 but they disassemble once the salt concentration rises above a

certain threshold.70 Hedrick et al. introduced urea functional groups,71 while Zhu

et al. introduced DNA base pairs72 into block copolymers to show that hydrogen

bonding can reduce micelle decomposition rates.

3.2.1.3 Approaches to Achieve Robust Intracellular Release

The chemical forces discussed above that make carriers retain drugs can conflict

with the need for a rapid and complete release at the target site. Drugs become

active only after liberation from their carriers.73,74 DOX that was stably bonded

to the nanoparticle core of poly(lactic-co-glycolic acid) (PLGA)75 or P(Asp)76


showed low or even no anticancer activity.77 The rate of drug release is also very

important because tumor cells have intrinsic and acquired drug-resistance

mechanisms to remove intracellular drugs,78,79 e.g. as a result of cell-membrane-associated multidrug resistance to efflux drugs79,80 and cell-specific drug

metabolism or detoxification.81 Tumor cells can also sequestrate some weakly

basic drugs in their lysosomes and use biomacromolecules to bind drugs to limit

their access to their targets. Thus, it is only the intracellular drugmolecules free to

bind to their targets that are useful therapeutically. Such free drug concentration

in the cytosol, herein referred to as the effective cytosolic drug concentration [D]

(effective [D] for short) determines the overall therapeutic efficacy.

Drug carriers that reach tumor cells are generally internalized by endocy-

tosis82,83 and routed to endosomes and then acidic lysosomes, as shown in

Figure 3.7. The internalized carrier can release the drug in one of two possible

ways or both: (1) within the lysosome, followed by drug diffusion, as illustrated

with the upper path in Figure 3.7, and (2) in the cytosol, following the carrier

escape from the lysosome, as illustrated in the lower path in Figure 3.7. For a

specific tumor cell, [D] is a function not only of the cellular uptake of the carrier

but also of its drug release rate (see Eq. 1 on the figure). If either ends up being‘‘too little, too late,’’ it can prevent reaching an effective [D].

(1) Intra-lysosome release

The intra-lysosome release mechanism (upper path in Figure 3.7) works for

most carriers that can be endocytosed into endosomes/lysosomes. The pH inendosomes decreases progressively, typically near 6 in early endosomes, near 5

in late endosomes, and about 4–5 in lysosomes.84 This acidic pH and the

Figure 3.7 Cytosolic drug accumulation by drug delivery: [D], effective drugconcentration in cytosol; Re, endocytosis rate of the carrier; RrL, drugrelease rate of the carrier in lysosomes; RrC, drug release rate of thecarrier in cytosol; RLDr, lysosomal drug release rate; RLCr, lysosomal-carrier escape rate; RR, the overall rate of drug removal by P-gp pumpsand drug consumption by other forms of drug resistance. Reprinted withpermission from ref. 8. Copyright 2012 Elsevier.

42 Chapter 3

special enzymes in lysosomes can trigger drug release from the carriers into

lysosomes.85 Because the harsh environment of lysosomes can easily degrade

drugs sensitive to acid or these enzymes,86,87 the drug must quickly diffuse out

into the cytosol to avoid deactivation.

Polymer–drug conjugates, in which the drugs are conjugated to the polymer

carriers via lysosomal pH-labile linkers, are the most popular design. Hydrazone

and cis-aconityl are examples of such a linker.88,89 Ulbrich et al. conjugated DOX

to N-(2-hydroxypropyl)methacrylamide (HPMA) copolymers via this hydro-

lytically labile spacer.90 The results showed a fast DOX release from the polymer

at intracellular pH 5, whereas at pH 7.4 the conjugates retained the drug.

Recently, they synthesized new biodegradable star conjugates consisting of

poly(amido amine) (PAMAM) dendrimer cores and HPMA grafts bearing DOX

via hydrazone bonds.89 The in vitro cytotoxicity and in vivo antitumor activity of

all such conjugates were higher than those of classic conjugates. Another example

is Wang et al.’s dual pH-responsive polymer–drug conjugate PPC-Hyd-DOX-

DA, which could respond to the tumor extracellular pH gradients via amide

bonds and the tumor intracellular pH gradients via hydrazone bonds.91

Lysosomal degradable peptides [e.g. glycylphenylalanylleucylglycine (GFLG)],

which are cleavable by lysosomal enzymes to release the drugs, are also used for

drug conjugation.74,92 For instance, DOX was conjugated to HPMA copolymers

via GFLG peptides to form a cleavable HPMA-GFLG-DOX conjugate.74

Lysosomal pH has also been used to trigger drug release from pH-sensitive

nanoparticles.93,94 For example, pH-sensitive micelles composed of reducible

poly(b-amino ester) (RPAE) cores dissociated rapidly in an acidic environment

and at high levels of reducing reagents, inducing fast intracellular release.94

Carriers with a core made from amine-containing hydrophobic polymers, such

as polyhistidine (PHis), can be protonated and thus dissolve in acidic lysosomes,

thereby releasing the drug.95 Our group showed that a rapid cytoplasmic release

from carriers could increase the anticancer activity of drugs.96,97

The additional advantage of such amine-containing polymers is that they

may also have endosomal membrane-disruption activity induced by a ‘‘proton

sponge’’ mechanism,98 and thus disrupt the lysosomal membrane and further

release the drug into the cytosol. Some specially designed polyacids, such as

poly(propylacrylic acid) (PPAA),99,100 were shown to disrupt endosomes at

pH 6.5 or below, causing the cytosolic release of cargo molecules.

(2) Intra-cytosol release

An alternative to the carriers designed for intra-lysosome release discussed

above is carriers designed for intra-cytosol release (lower path in Figure 3.7). Such

intra-cytosol-release carriers retain the drug until escape from the endosome/

lysosome101 and then release the drugs into the cytosol, hence avoiding lysosomal

drug retention and degradation. This is particularly important in small interfering

RNA (siRNA) or gene delivery and thus various approaches have been explored

to facilitate the endosomal release of DNA or RNA complexes.102,103 In this

approach the carriers must respond to the lysosomal environment for lysosomal

escape and to the cytosolic environment for drug release.


Stealth carriers, such as HPMA104 and pegylated particles, cannot diffuse

through the lysosomal membrane and thus can be retained in the lysosomes for

a long time. For instance, PEG-PCL particles were found confined in

lysosomes.105 Thus, they must be functionalized with lysosomal membrane-

destabilizing polymers such as PPAA,99,100 pH-dependent fusogenic pep-

tides,106,107 or cationic polymers such as polyethylenimine (PEI)108 or

histidine-rich peptides or polymers.109 For cationic polymers or peptides, on

the other hand, it is important first to mask their cationic charges (from

primary and secondary amines) at physiological pH, so the carriers can be used

for i.v. administration. However, once inside the tumor lysosome, the cationic

charges are recovered to lyze the lysosomal membrane for escape. Such a

‘‘negative-to-positive charge-reversal’’ method makes the carrier stealthy in

circulation, but enables endosomal lysis, once in lysosomes.108,110

Removal of a cleavable PEG layer can also allow lysosomal escape.111 For

instance, a PEG-cleavable lipid, via an acid-labile vinyl ether linker, was used

for pegylation of (1,2-dioleoyl-sn-glycero-3-phosphoethanolamine) (DOPE)

liposomes. At acidic lysosomal pH the vinyl ether linker hydrolyzed and the

PEG layer was removed from the DOPE liposomes, enabling DOPE, which

has excellent fusogenic capacity, to fuse with the lysosomal membrane for

escape.112 Disulfide linkages were also used to detach PEG and make the drug-

loaded carriers quickly escape from the endosomes.113 After the particles were

internalized by cells and trapped by endosomes, the PEG layer was removed.

The exposed particles interacted with the endosomal membrane, increased

the endosomal pressure, or both, resulting in destruction of the endosomal

membrane to enable effective endosomal escape.113

Most carriers reaching the cytoplasm have already experienced an initial

burst release and are in a slow, diffusion-controlled drug release process, e.g.

nanoparticles with cores made of solid glassy polymers such as PCL or

polylactide (PLLA).114 According to Eq. 1 (see Figure 3.7), such a slow drug

release profile may not be able to lead to a high [D] lethal to cancer cells. Thus,

carriers responding to cytosolic signals have been developed for faster drug

release. The most common is a cytosolic redox signal resulting from an

elevated intracellular GSH concentration (y10 mM) compared to that in the

bloodstream (y2 mM).115 GSH can effectively cleave the disulfide bonds to

release conjugated drugs.66,116 It is thus used to trigger decomposition of

micelles with hydrophobic parts linked by disulfide bonds117 or other carriers

crosslinked118 or gated119,120 with disulfide linkers. It has also been observed

that removal of the PEG corona could increase the drug release rate.121,122

3.2.2 2S: Stealthy in Circulation and Tumor Penetration versusSticky to Tumor Cells

The second major material challenge is how to impart nanocarriers’ stealth

ability to circulate in blood for a long time and after extravasation to penetrate

44 Chapter 3

deep into the tumor, to the cells away from the blood vessels, but become

effectively sticky upon interacting with tumor cells for fast cell internalization.

To be stealthy for a long circulation time in the blood compartments hasbeen recognized as essential for a nanocarrier to achieve passive tumor

targeting,123,124 whereas transport in the tumor tissue after extravasation has

been gradually realized in recent years.125,126 Tumor resistance to anticancer

drugs not only involves the cellular and genetic drug resistance mechan-

isms,78,127 but also the physiological barriers of solid tumor tissues.128,129 It is

found that tumor drug distribution is not uniform. Drugs are rich in the areas

surrounding the blood vessels and the concentration declines sharply away

from the blood vessels, owing to the compact structure of tumor tissues.130

Thus, the most aggressive tumor cells located in these hostile microenviron-

ments (low pH and low pO2) are actually exposed to few drugs.7 Moreover, the

exposure of those cancer cells to sublethal concentrations of anticancer drugs

may facilitate the development of resistance.7 Therefore, it is important for the

nanocarrier to remain stealthy after extravasation for tumor penetration. It

can be imagined that a nanocarrier strongly interacting with surrounding cells

and matrix will be trapped there and cannot travel a long distance.

3.2.2.1 Approaches to Stealth Surfaces

(1) In circulation

The nanocarrier’s stealth character hinges on many factors, including

surface properties,123 size,131 and even shape.132,133 In circulation, those with

molecular weights below the renal threshold (e.g. 40 kDa for PEG) or sizesbelow 5 nm are rapidly cleared from the blood by glomerular filtration,86 while

those with diameters above 200 nm will be scavenged by the RES, mainly the

liver and spleen.131,134

Most stealth carriers capable of avoiding opsonization135 and interaction

with the mononuclear phagocyte system (MPS)131 are made from

HPMA,136,137 PEG, or polysaccharides138 (e.g. heparin139). Nanoparticles

coated with a layer of these polymers become stealthy by both hydration and

steric hindrance.140 For example, pegylation of particles or liposomes is well

established,135,141 and the DOX-loaded stealth liposome named Doxil1 was

approved by the FDA for cancer therapy.142 Huang et al. reported that, on

100 nm liposomes pegylated with 1,2-distearoyl-sn-glycero-3-phosphoethanola-mine-PEG2000 (DSPE-PEG2000), PEG chains were arranged in a mushroom

configuration at a DSPE-PEG fraction less than 4 mol% but in a brush

configuration at a DSPE-PEG content greater than 8 mol%.124 The high density

of PEG chains on the liposome surface with the brush configuration was the key

to reduce liposome liver sequestration.124 Discher et al. incorporated the PEG

brushes onto polymersomes and obtained polymersomes having a blood

circulation time two-fold longer than pegylated liposomes.143 Dai et al.

pegylated single-wall carbon nanotubes (SWNT) and found that, with theincrease of linear PEG chain length from 2 kDa to 5 kDa, the blood circulation


time of pegylated SWNTs was significantly extended, but a further increase of

the PEG chain length showed no significant effect.144 Although pegylation

reduces the recognition of the carriers by the MPS system and thereby extends

their blood circulation time, the ‘‘accelerated blood clearance (ABC)’’

phenomenon was observed upon repeated injection of pegylated liposomes145,146

due to IgMbound to pegylated liposomes secreted into the bloodstream after the

first dose.147 Such an immune reaction against the pegylated liposomes occurred

in the spleen at least 2–3 days after the first administration.145,146

The carrier shape is also recognized as an important parameter that can

substantially affect the blood circulation time. In fact, Mitragotri et al. reported

that the particle shape, not size, played a dominant role in phagocytosis of

polystyrene (PS) particles of various sizes and shapes: the rod-like particles

entered the cells much faster.148 Discher et al. found that flexible worm-like

micelles efficiently evaded the RES and circulated in the blood for a week,149,150

much longer than spherical micelles. Dai et al. found that carbon nanotubes

pegylated with long PEG chains exhibited a long blood circulation time (t1/2 5

22.1 h) upon intravenous injection into mice.151 All these studies suggest that

particle phagocytosis can be inhibited by minimizing its size-normalized

curvature.148,152 Thus, particle shape is an important variable to make it remain

stealthy in circulation long enough for enhanced tumor accumulation.149,150,153

(2) In tumor tissue

Solid tumors are characteristic of poorly structured blood vessels,154 a stiff

extracellular matrix (ECM),155–157 tightly packed cells,158 high interstitial fluid

pressure (IFP),159,160 and drug metabolism and binding.126 Together they

impose strong diffusion barriers to nanocarriers, and even small molecules

(Figure 3.8).125,161

For instance, it seems intuitive that as long as a nanocarrier extravasates

from tumor blood capillaries and releases the carried small molecular drugs,

the drug molecules will diffuse deep into the tumor tissue. Actually, free drugs,

either hydrophobic or carrying positive charges, cannot migrate far from the

nanocarrier due to their avid binding.162 Diffusion of larger macromolecules,

such as bovine serum albumin (BSA, 68 kDa, 9 nm in diameter) and

immunoglobulin G (IgG, 150 kDa, 11 nm in diameter), in the tumor ECM

is also hindered compared to that in buffered saline.163 After extravasation,

dextrans with a molecular weight between 40 and 70 kDa (and a diameter of

11.2–14.6 nm) were observed to be concentrated near the vascular surface.130

Apparently, nanocarriers, which are larger in size than BSA and dextrans, are

likely to face greater difficulties in tumor penetration.161 Chan et al.

systematically examined the effect of nanoparticle size on tumor penetration

using sub-100 nm pegylated gold nanoparticles.164 As expected, larger

nanoparticles appeared to stay near the vasculature while smaller nanopar-

ticles (20 nm in diameter) rapidly diffused into the tumor matrix (Figure 3.9).

Similarly, Lee et al. demonstrated that PEG-PCL micelles with a mean

diameter of 25 nm diffused further away from the blood vessels compared to

those with diameters of 60 nm, which mainly remained in the perivascular

46 Chapter 3

regions.165 Furthermore, Kataoka et al.166 and Liang et al.167 also proved that

small-sized nanocarriers were essential for improved diffusion.168 However,

small-sized nanocarriers possess a high probability to be fast cleared during the

circulation, as mentioned above.

Figure 3.8 Scheme of solid tumor tissue which is characteristic of stiff ECM andcompact tumor cells. Adapted with permission from ref.125. Copyright2012 Elsevier.

Figure 3.9 Size-dependent penetration of nanoparticles within tumor tissues.Reprinted with permission from ref. 164. Copyright 2009 AmericanChemical Society.


Besides size, the surface charge of nanocarriers also influences their

penetration into tumor tissues. Cationic liposomes (150 nm) were observed not

able to travel far into the tumor interstitium.169 Recently, Forbes et al. compared

the penetration of oppositely charged gold nanoparticles (+30 vs. 236 mV, 6 nm)

into cylindroidal cell aggregates.170 Cationic nanoparticles were taken up by the

proliferating cells on the periphery of the cylindroids, whereas anionic

nanoparticles were better at penetrating the extracellular matrix and entered

hypoxic necrotic cells in the core of the mass. As a matter of fact, the extracellular

matrix presents as an effective electrostatic bandpass, suppressing the diffusive

motion of both positively and negatively charged objects, which allows

uncharged particles to easily diffuse through while effectively trapping charged

particles (Figure 3.10).171 Jain et al. demonstrated that the optimal particles for

delivery to tumors should be neutral after exiting the blood vessels.172

Another issue that needs addressing is affinity.173 The affinity plays an

important role in antibody-based tumor targeting nanocarriers. It was visualized

that the antibody distributed mostly in perivascular regions rather than homo-

geneously in tumor cells.174 Reports revealed there was an inverse relationship

between affinity and penetration, i.e. the antigen–antibody interaction in the

tumor tissue imposed a binding-site barrier that retarded antibody penetration

and caused a heterogeneous distribution.175–177 The higher the affinity of binding

and the higher antigen density caused fewer free molecules to be able to penetrate

farther into the tumor interstitium.175,176 Increasing the antibody dose gave better

penetration and more uniform distribution.175

Therefore, it is clear that to deliver a sufficient drug concentration to the

tumor center region lacking vascular perfusion, where the most aggressive and

resistant cells reside, the nanocarrier should not release the carried drug after

Figure 3.10 Scheme of the ECM exerts the filtering function in tumor tissue.Charged particles (red, blue) are trapped in the respective region ofopposite charge (blue, red), while neutral particles (gray) can diffusenearly unhindered.

48 Chapter 3

extravasation but should further diffuse deep into the tumor. This requires the

nanocarrier to remain slippery and have as small a size as possible. Thus, it is

better for the nanocarrier to be neutral and not to present any binding groups

(including targeting groups) until reaching the center of the tumor.

3.2.2.2 Approaches to Becoming Sticky to Tumor Cells forCellular Uptake

After reaching the targeted region the nanocarrier should efficiently enter the

cells for drug release. Now the same properties that impart stealth to the

nanocarrier prevent it from cellular uptake by tumor cells. Nanocarriers that are

negatively charged will be repelled from the cell membrane due to the electrostatic

repulsion. The PEG corona of pegylated polymeric micelles or liposomes retards

their interaction with cell membranes due to steric hindrance. Thus, once in the

tumor, the carrier must become cell-binding or sticky to targeting tumor cells for

fast cellular uptake.178 The challenge is how to reconcile these two opposite

requirements, stealth for circulation and diffusion versus sticky for targeting. For

instance, it is well known that positively charged carriers reliably stick to cell

membranes due to electrostatic adsorption triggering fast cellular uptake, but

positively charged carriers are not suitable for in vivo applications because they

are systemically toxic179 and have short circulation times.180

One strategy to convert a carrier from stealth circulation to sticky targeting

is to equip it with PEG groups that are cleavable upon encountering a tumor-

specific stimulus. Once the PEG chains are removed, the bare particle can be

adsorbed onto the cell membrane. Toward this end, Thompson et al. prepared

acid-labile PEG-conjugated vinyl ether lipids to stabilize fusogenic DOPE

liposomes.181 At lower pH, the PEG layer was removed by the acid-catalyzed

hydrolysis of the vinyl ether bond, triggering membrane fusion. Similarly,

Harashima et al. connected PEG to the lipid through a matrix of

metalloproteinase (MMP)-cleavable peptide.182,183 MMP is overexpressed in

tumor-tissue angiogenesis, invasion, and metastasis184 and thus the peptide can

be degraded quickly in tumors. They prepared a multifunctional envelope-type

nano device (MEND) using the PEG-peptide lipid and found that pDNA

expression was dependent on the MMP expression level in the host cell.

Positive charges can promote carrier adsorption on the negatively charged

membrane and hence trigger adsorption-mediated endocytosis. Thus, an

alternative is to use tumor extracellular acidity to impart positive charges to the

carrier by a ‘‘charge-reversal’’ technique (illustrated in Figure 3.11). Amine-

containing carriers, such as PCL-b-PEI,108 poly(L-lysine) (PLL),185 and PAMAM

dendrimers,110 were amidized to acid-labile b-carboxylic acid amides to make

them negatively charged at physiological pH. Once in weakly acidic tumor

extracellular fluid, the amides hydrolyzed and regenerated the amines with

cationic charges, which led to fast cellular uptake (Figure 3.11A). In yet another

example, a pH-responsive layer becomes positively charged at tumor extracellular

acidity but collapses, forming a middle layer, at neutral pH (Figure 3.11B).16 Bae


et al. reported tumor extracellular pH-triggered TAT-presenting micelles. The

TAT moieties were anchored to a PEG micelle corona and shielded at pH .7.0

by their electrostatic complexation with poly(methacryloyl sulfadimethoxine)

(anionic PSD)-PEG (PSD-b-PEG) diblock copolymer. At pH 6.6, however, PSD

turned to a nonionized form and fell off the TAT, exposing it and enabling the

micelle a fast cellular uptake.186 Another design was to anchor TAT onto the PEG

corona through a pH-sensitive PHis spacer. At pH 7.4, the PHis was water

insoluble, which kept the TAT moieties buried in the PEG corona. At pH lower

than 7.2, however, ionization of the PHis spacer made it water soluble, which

stretched it, exposing the TAT on the corona surface.187

The most common approach enabling a carrier to become sticky to the cell

membrane is to decorate it with a ligand whose receptors are overexpressed on the

cancer-cell membrane. The ligand–receptor binding enables receptor-mediated

endocytosis, promoting cellular uptake.11,188 Only a few ligands are needed for

rapid internalization.189More ligand groups can theoretically increase uptake, but

a high surface ligand density may make the carrier less stealthy as a result of

opsonization-mediated clearance.190 Many examples of targeting ligands include

folic acid,191 peptides,192–194 antibodies,195–197 transferrin,198 aptamers,199,200 and

other moieties201 that have been tested and subsequently reviewed.2,202

Figure 3.11 (A) The charge-reversal concept for drug delivery. Reprinted withpermission from ref. 14. Copyright 2010 Elsevier; (B) The pH-responsivethree-layered nanoparticles (3LNPs). Reprinted with permission fromref. 16. Copyright 2008 American Institute of Chemical Engineers.

50 Chapter 3

3.3 The Material Excipientability and ProductionProcess Scale-Up Ability

The 2R2S capability for nanocarriers discussed in the previous sections

determines the adsorption, distribution, metabolism, and excretion (ADME)

of the carried drug. Such a nanocarrier simultaneously having 2R2S capability

can deliver a high cytosolic drug concentration and give rise to high

therapeutic efficacy. However, this is not sufficient for it to be trans-

lational.203–205 The nanocarrier itself should also have proper ADME.

According to Choi and Frangioni, safety and clearance (renal or hepatic)

and a proper stealth surface should be included among the basic criteria for

clinical translation of formulation/materials administered to humans,205 ‘‘from

the benchtop to the bedside’’ translation. Thus, a nanocarrier must meet the

requirements for the pharmaceutical excipient for i.v. uses. For simplicity, this

ability of the nanocarrier material(s) to be used or approved to be an excipient,

herein denoted as excipientability, is the second element for a nanocarrier to be

translational (see Figure 3.2). It goes without saying that the production of the

nanocarrier and the resulting nanomedicine should be able to be scaled-up and

establish the required GMP, or scale-up ability, for short. Some of these

important points of the two key elements are summarized as follows.

1) Safety. The nanocarrier itself should have proper ADME and no

nanotoxicity, and should be nontoxic and easy to excrete completely from the

body via the liver (into bile) or the kidneys (into urine) or both. This is because

retention of polymers or nanosized materials in the body, even inert polymers

like polyvinylpyrrolidone (PVP),206–208 can cause health problems. The

threshold for rapid renal excretion is about 5.5 nm in hydrodynamic diameter.

This corresponds to a molecular weight of about y 45 kDa for HPMA209 and

40 kDa for PEG.86

2) Approval. In order to expedite and increase the probability of the

approval success, the carrier should have a clear and simple structure with

known degradation products. An even better case would be that it is made of

FDA-approved building blocks.

3) Production scale-up. This involves the feasibility of making large volumes

of consistently reproducible quality to establish GMP. For instance, because

the molecular weight of a polymer–drug conjugate strongly affects its

pharmacokinetics, the polymer itself must have consistently low polydispersity

and reproducible average molecular weight from batch to batch. The same

applies to drug-loaded micelles made of block copolymers, such as PEG-PCL,

in addition to reproducible particle size, particle-size distribution, and drug-

loading efficiency and content. As the micelle structure becomes more and

more complicated, the number of quality control parameters drastically

increases,14,210 which makes it more and more difficult to produce an

acceptably consistent formulation. Also, although not crucial to clinical

success, it is also worth considering a high, ideally close to 100%, drug-loading


efficiency to simplify the manufacturing process and minimize losses of these

very expensive anticancer drugs.

4) High drug-loading content. In current commercial formulations, thedrug-loading content tends to be on the low side.211–213 High drug-loading

contents are needed to minimize the body’s exposure to excipient carrier

matter, even if it is biocompatible and relatively benign. For instance, PEG-

containing liposomal carriers may induce acute immune toxicity, manifested in

hypersensitivity reactions (HSRs).214,215

3.4 Challenges of Rational Design for TranslationalNanomedicine

With the above analysis in mind, it is clear that the key to translational

nanomedicine is to develop nanocarriers with optimal 2R2S capability,excipientability, and scale-up ability.

As for the nanocarrier 2R2S capability, we still do not have ones that can

fully and simultaneously achieve the 2R2S capability, despite a large volume ofthe scientific literature on each topic separately, or on various subsets of them,

giving rise to unsatisfied therapeutic efficacy and side effects. As a

consequence, a particular problem of those systems is that a large majority

doses of the drugs are still sequestrated in the liver or spleen, even though the

tumor drug accumulations are indeed enhanced compared to free drugs.133,216

For instance, the PF-PTX micelles217 and IT-101 CPT conjugates218 give drug

accumulation in tumors much better than Taxol1 and CPT, respectively, but

the total amounts of drugs accumulated in the liver were still about 4.5 and 3.5times of those in tumors. In many cases, only a few percent of the injected

drugs were in the tumors. Thus, for many nanomedicine systems, liver toxicity

is the killer for further developments. Other necessities are how to achieve

effective cellular uptake of the nanocarriers once in the tumor and robust

intracellular release. Delayed or insufficient intracellular release directly leads

to lower cytotoxicity than the free drugs.219,220

The material excipientability of nanocarriers and the production scale-up

ability of the nanocarriers and their nanomedicine systems are equally

important. For instance, a large variety of inorganic nanomaterials and

sophisticated polymeric nanostructures have been proposed and investigated

as nanocarriers for cancer drug delivery. These studies provide useful proof-of-concepts and rich insights into various aspects of cancer drug delivery essential

to the design of nanocarriers towards 2R2S capability, but those aimed at

clinical applications must comprehensively design and characterize their

materials, nanosize effects, and scale-up ability. Of the three, the material is

the basic concern for a translational nanocarrier. If the material used for the

nanocarrier is not proper for in vivo clinical uses (for instance, inherently toxic

or non-clearable from the body), the resulting nanocarrier, even with perfect

nanosize effects and 2R2S capability, would not be able, or take animpractically long time, to be translated into clinics. Thus, except for proof-

52 Chapter 3

of-concepts, it is better to look into these issues early at the bench in order for a

successful nanocarrier to move forward quickly.

3.5 Conclusion

The challenge to develop truly translational nanocarriers and nanomedicine is

to use excipientable materials and processes of scale-up ability to produce

nanocarriers with optimal 2R2S capability. While research aimed at proof-of-

concepts remains important, it is important to increasingly focus on

comprehensive approaches or systems that include all the three key elements

as early as possible in the innovation chain, to speed up developments of

translational nanomedicine.

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CHAPTER 3 Rational Design of Translational Nanocarriers · Rational Design of Translational Nanocarriers 33. ... Using a stepwise pH-controlled process, ... (PDEA) middle layer, and